The quantitative determination of analytes in biological fluids is useful in the diagnosis and treatment of physiological abnormalities. For example, determining the glucose level in biological fluids, such as blood, is important to diabetic individuals who must frequently check their blood glucose level to regulate their diets and/or medication.
Electrochemical methods have been used for such purposes.
An electrochemical biosensor may use an analyte specific enzyme, such as glucose oxidase or glucose dehydrogenase, to catalyze the oxidation of glucose in a whole blood sample. During the catalytic oxidation by the enzyme, the redox center of the enzyme accepts the electrons from the analyte.
This redox center could be the flavin adenine dinucleotide (FAD) of glucose oxidase, or the enzyme's cofactor such as pyrroloquinoline quinone (PQQ) for the glucose dehydrogenase. The electrons acquired by the enzyme then may be moved to the electrode by a mediator, which is converted to a reduced form through oxidation of the enzyme. Finally, the reduced form of the mediator, such as the ferrocyanide species of the ferricyanide/ferrocyanide redox pair, is oxidized at the electrode to generate a measurable current.
This process may be represented by the following equations:Glucose+EOx==ERed+Product  (1)ERed+nMedOx==nMedRed+EOx  (2)MedRed==MedOx+ne−  (3)where EOx and ERed are the oxidized and reduced forms of the redox center of the enzyme, respectively, while MedOx and MedRed are the oxidized and reduced forms of the mediator, respectively. The product of the enzymatic reaction may be gluconic acid or gluconolactone.
One electrochemical method, which has been used to quantify analytes in biological fluids, is coulometry. For example, Heller et al. described the coulometric method for whole blood glucose measurements in U.S. Pat. No. 6,120,676. In coulometry, the analyte (glucose) concentration is quantified by exhaustively oxidizing the analyte within a small volume and integrating the current over the time of oxidation to produce the electrical charge representing the analyte concentration. In other words, coulometry captures the total amount of glucose within the sensor strip.
An important aspect of coulometry is that towards the end of the integration curve of charge vs. time, the rate at which the charge changes becomes relatively constant to yield a steady-state condition. This steady-state portion of the coulometric curve forms a relatively flat plateau region in the curve, thus allowing accurate determination of the corresponding current. However, the coulometric method requires the complete conversion of the entire volume of analyte. As a result, this method is time consuming and does not provide the fast results which users of electrochemical devices, such as glucose-monitoring products, demand. Another problem with coulometry is that the small volume of the sensor cell must be controlled in order to provide accurate results, which can be difficult with a mass produced device.
Another electrochemical method which has been used to quantify analytes in biological fluids is amperometry. In amperometry, current is measured at the end of a period at a constant potential (voltage) across the working and counter electrodes of the sensor strip. The current is used to quantify the analyte in the biological sample. Amperometry measures the rate at which the electrochemically active species, and thus the analyte, is being oxidized or reduced. Many variations of the amperometric method for biosensors have been described, for example in U.S. Pat. Nos. 5,620,579; 5,653,863; 6,153,069; and 6,413,411. The amperometric method samples the analyte concentration near the electrode surface by measuring the current that is proportional to the diffusion rate and the bulk concentration of the analyte.
A disadvantage of the amperometric method is the non-steady-state nature of the current after applying a potential. The rate of current change with respect to time is very fast initially and becomes slower as the analysis proceeds due to the changing nature of the underlying diffusion process. Until the consumption rate of the reduced mediator at the electrode surface equals the diffusion rate, a steady-state current cannot be obtained. Thus, measuring a current during a non-steady-state time period may be associated with more inaccuracy than a measurement taken at a steady-state time period.
One important aspect of measuring analytes in whole blood samples is the effect of hematocrit. Hematocrit is the volume of red blood cells (RBC) expressed as a percentage of the volume of RBC in a whole blood sample. The hematocrit value for whole blood samples ranges from about 20 to 60% and is typically about 40%.
Reagent biosensors include any system that can detect glucose in a blood specimen via an electrochemical reaction. Examples of reagent biosensors include Ascensia AUTODISC® and Elite® biosensors available from Bayer HealthCare, LLC of Elkhart, Ind.; Precision® biosensors available from Abbott in Abbott Park, Ill.; Accucheck® biosensors available from Roche in Indianapolis, Ind.; and OneTouch Ultra® biosensors available from Lifescan in Milpitas, Calif.
Typical electrochemical sensor strips contain a working electrode, a counter electrode, and an optional third electrode. A reference potential may be provided to the system by the counter electrode, if configured appropriately, or by the optional third electrode. A reagent layer with an enzyme such as glucose oxidase or glucose dehydrogenase and a mediator such as ferricyanide or ruthenium hexaamine is printed or deposited onto the working electrode or onto the working and counter electrodes with a polymer as the binder.
Examples of polymers used as the binder of the reagents include CMC (carboxyl methyl cellulose) and PEO (polyethylene oxide). The addition of various types and molecular weights of polymers to the reagent layer may assist in filtering red blood cells, preventing them from coating the electrode surface.
Preferably, the sensor strip is made by printing electrodes on an insulating substrate using multiple techniques, such as those described in U.S. Pat. Nos. 6,531,040; 5,798,031; and 5,120,420. The reagent can be co-printed onto the working and counter electrodes with a mixture of a glucose oxidizing enzyme such as glucose oxidase, a mediator such as ferricyanide, a hydrophilic polymer such as polyethylene oxide (PEO) and an appropriate buffer, such as a citrate buffer.
Alternatively, a different reagent chemistry can be either printed or micro-deposited separately onto the working and counter electrodes using the method described in a U.S. provisional patent application filed Oct. 24, 2003, Ser. No. 60/513,817 with the reagent on the working electrode containing the enzyme, the mediator, the polymer and that on the counter electrode containing a soluble redox species, which could be the same as the mediator or different, and a polymer. In one embodiment, the polymer used in micro-deposition is carboxyl methyl cellulose.
Examples of suitable bench-top electrochemical instruments which may be used for reading reagent biosensors according to the present invention include, but are not limited to, the BAS 100B Analyzer available from BAS Instruments in West Lafayette, Ind.; the CH Instrument Analyzer available from CH Instruments in Austin, Tex.; the Cypress Electrochemical Workstation available from Cypress Systems in Lawrence, Kans.; and the EG&G Electrochemical Instrument available from Princeton Research Instruments in Princeton, N.J. Examples of portable instruments include the Ascensia Breeze® and Elite® meters of Bayer Corporation.
A biosensor for glucose may have an enzyme and a mediator deposited on the electrodes. The ability of this sensor to measure glucose is affected as the RBC block the diffusion of the relevant reagents within the blood sample. Since the amperometric current is directly proportional to the diffusion of the reduced form of the mediator, the hematocrit will have a significant impact on the accuracy of the glucose measurements. Depending on the hematocrit level in a whole blood sample, the RBC cause a bias in the glucose readings.
Various methods and techniques have been proposed in an attempt to reduce the hematocrit effect of the whole blood on the resulting glucose measurements. For example, Ohara et al. in U.S. Pat. No. 6,475,372 disclosed a method of using the ratio of currents from a forward and a reverse potential pulse to compensate for the hematocrit effect in electrochemical glucose measurements. McAleer et al. in U.S. Pat. Nos. 5,708,247 and 5,951,836 disclosed a reagent formulation using silica particles to filter the RBC from the electrode surface, thus reducing the hematocrit effect. Carter et al. in U.S. Pat. No. 5,628,890 disclosed a method of using a wide spacing of the electrodes combined with mesh layers to distribute the blood sample for the hematocrit effect.
These conventional techniques for reducing the bias attributable to the hematocrit effect include (a) co-deposition of a polymer to minimize the hematocrit effect, (b) addition of various kinds of fused silica to enforce the filter effect for the polymer layer, (c) compensation coefficients based on the ratio of currents from a forward and a reverse potential pulse, and (d) self-compensation by utilizing the existing solution resistance of the whole blood samples. Although these methods may be useful, conventional glucose sensors continue to exhibit significant analytical bias attributable to the hematocrit effect. Thus, it would be desirable to provide systems for quantifying analytes in biological fluids, in particular the glucose content of whole blood, which reduces bias from the hematocrit effect.